This invention relates magnetic resonance imaging (MRI) systems, to magnet systems for producing a homogeneous imaging field for MRI and, particularly, to open magnet systems having two orthogonal polarization axes and providing a remote region of field homogeneity along with planar gradient coils for delivering gradient fields for spatial encoding in a remote target field region. More particularly, this invention relates to MRI magnet systems having a pair of current loops wound about a substantially planar ferromagnetic core, the core preferably including an orthogonal end piece.
There are known whole body MRI magnets (super-conductive, resistive iron core magnets, and permanent magnets), which produce the background Bo field used in MRI. The useable imaging volume in these magnets is in the region where the field is an extremum and provides a region of substantial field homogeneity. This volume is located in the air space centrally located between field sources. Thus, typically, MRI magnets are designed to provide a homogeneous magnetic field in an internal region within the magnet, e.g., in the air space of a large central bore of a solenoid or in the air gap between the magnetic poles of a C-type magnet. A patient or object to be imaged is positioned in the homogeneous field region located in such air space. In addition to the main or primary magnet that provides the background magnetic field Bo, the MRI system typically has gradient and rf coils, which are, used respectively for spatial encoding and exciting/detecting the nuclei for imaging. These gradient field and rf coils are typically located external to the patient inside the geometry of the Bo magnet surrounding the central air space.
Prior art electromagnets such as described by Watson et al and Muller et al. and other prior art iron core magnets typically have a structural design to provide a maximum magnetic field strength at a large central air space. In addition, those types of the prior art magnets, of the iron core electro- or permanent type, have a substantial edge fringe field effect, which makes it difficult to image beginning immediately at the magnet edge or even proximal to the edge of the magnet due to lack of sufficient field homogeneity.
In U.S. Pat. No. 5,049,848 a magnet configuration for MRI mammography is disclosed. The magnetic structure 50 has a rectangular shaped magnet with at least two parallel magnetic source 5,6 connected by a ferromagnetic core flux path defining an air gap for imaging. A remote shimming C-shaped magnetic source is preferably used for shimming to decrease the front edge fringe effect of the magnetic structure 50 to create a relatively homogeneous field in the air gap beginning at the front edge for effective imaging.
Solenoidal MRI magnets (superconductive, resistive) as well as iron core C and E shape electromagnets or permanent magnets are known for imaging of the whole body and its extremities. However, such whole body MRI magnets are not generally well-suited for treatment of the patient with other modalities or for minimally invasive surgical procedures guided by real time MRI because of the limited access of the surgeon to the patient. This limited access results from the field producing means surrounding the imaging volume. Electromagnets of the C or E type iron core configuration have been designed to offer a partially open access to the patient, however, the access is still very limited with typical air gaps of only 40 cm between the pole pieces of a C type magnet. U.S. Pat. No. 5,378,988 describes a MRI system, which can provide access for a surgeon or other medical personnel, using a plurality of C-shape solenoidal magnets oriented to form an imaging volume in a central region of the magnets.
Another type of magnet specifically designed for interventional surgical guidance is General Electric""s Magnetic Resonance Therapy device, which consists of two superconducting coils in a Helmholtz coil type arrangement (see U.S. Pat. No. 5,677,630). The air gap for this commercial magnet is about 58 cm, which typically permits access by one surgeon.
None of those prior art magnets or MRI systems is ideal with regard to simultaneously offering real time imaging and fully open access to the patient. Many surgical procedures require multiple surgeons together with an array of supporting equipment and, thus, a fully open magnet configuration for a MRI system for interventional procedures is desirable. In addition, such open magnet configuration is desirable for patients that have claustrophobia.
Applications other than MRI have used magnets that produce a useful field region outside the magnet geometry. U.S. Pat. No. 4,350,955 describes means for producing a cylindrically symmetric annular volume of a homogeneous magnetic field remote from the source of the field. Two equal field sources are arranged axially so that the axial components of the fields from the two sources are opposed, producing a region near and in the plane perpendicular to the axis and midway between the sources where the radial component of the field goes through a maximum. A region of relative homogeneity of the radial component of the background field Br may be found near the maximum. The large radial field is generally denoted as the Bo background field in MRI applications. See also, J. Mag. Resonance 1980, 41:400-5; J. Mag. Resonance 1980, 41:406-10; J. Mag. Resonance 1980, 41:411-21. Thus, two coils producing magnetic fields having opposing direction are positioned axially in a spaced relationship to produce a relatively homogeneous toroidal magnet field region in a plane between the magnets and perpendicular to the axis of cylindrical symmetry. This technology has been used to provide spectroscopic information for oil well logging but has not been used for imaging.
U.S. Pat. No. 5,572,132 describes a magnetic resonance imaging (MRI) probe having an external background magnetic field Bo. The probe has a primary magnet having a longitudinal axis and an external surface extending in the axial direction and a rf coil surrounding and proximal to the surface. The magnet provides a longitudinal axially directed field component Bx having an external region of substantial homogeneity proximal to the surface. Comparing this magnet geometry to that of U.S. Pat. No. 4,350,955, it has a background Bo field with a cylindrically symmetrical region of homogeneity. However, this magnet described in the copending application provides such a field in the axial direction (i.e., longitudinal axis direction) whereas the other provides a background Bo field in the radial or r direction (i.e., radial direction). Preferably, the Bo field is provided by two magnets spaced axially and in axial alignment in the same orientation and wherein said region of homogeneity intersects a plane that is located between the magnets and that is perpendicular to the axis. For MR imaging, surrounding the primary magnet are r-, z- and xcfx86-gradient coils to provide spatial encoding fields.
It is desirable to have new and better devices and techniques for biomedical MRI applications such as open magnet MRI systems for imaging while performing surgery or other treatments on patients or for imaging patients that have claustrophobia. It is also desirable to have portable devices and imaging techniques that could be applied to a wide variety of imaging uses.
U.S. Pat. No. 5,744,960 describes a planar MRI system having an open magnet configuration comprising two pairs of planar pole pieces that produces a magnetic field having a substantial remote region of homogeneity.
U.S. Pat. No. 5,914,600 describes an open solenoidal magnet configuration comprises a pair of primary solenoidal coils and, located within the primary coil geometry, a bias coil system, the coils emitting an additive flux in the imaging region to generate a resulting field which provides a remote region of substantial field homogeneity.
U.S. Pat. No. 6,002,255 describes an open, planar MRI system having an open magnet configuration including a planar active shimming coil array that produces a magnetic field having a substantial remote region of homogeneity. The MRI system also includes spatial encoding gradient coils and a rf coil, each preferably having a planar configuration.
The prior art magnet configurations that provide a primary background magnetic field having a remote region of substantial field homogeneity typically comprise a primary magnet system having spaced primary field emission surfaces and, located between the spaced field emission surfaces, a bias magnet system having spaced bias field emission surfaces that emit an additive flux in the imaging region to generate a resulting field which provides a remote region of substantial field homogeneity. The spaced primary field emission surfaces typically are the pole pieces of a primary magnet or a solenoidal magnet facing the target region.
However, a fundamental problem for obtaining an open magnet having maximal accessibility for a surgeon to conduct MRI guided surgery results from the fact that conventional magnet systems exhibit a substantial drop in magnet efficiency when providing an open volume of magnetic field that is large enough for surgery to be conducted therein. Conventional iron core C-type magnet configurations provide a target field volume between coaxial pole pieces. That type of magnet configuration, even with air gap enlargement (reducing magnet efficiency), still has limited accessibility for MRI guided surgery. MRI magnet systems in the form of one side xe2x80x9cpancake typexe2x80x9d magnet configurations (e.g. U.S. Pat. No. 5,331,282) generally have a set of coaxial circular coils with alternating polarity and axially shifted positions and provide relatively low level of remoteness and require a large diameter magnet for adequate field strength, thus, inhibiting accessibility to the region of field homogenity. So-called xe2x80x9copenxe2x80x9d solenoid superconductive magnets (e.g., U.S. Pat. No. 5,677,630) provide better accessibility and larger field of view (FOV) but accessibility still is limited by axial distance between two solenoidal magnets (which typically is about the width of a person""s shoulders). The planar open magnet systems mentioned above (e.g., U.S. Pat. No. 5,378,988; U.S. Pat. No. 5,744,960; U.S. Pat. No. 5,914,600 and U.S. Pat. No. 6,002,255) provide complete openess for excellent accessibility but suffer still from a limitation in magnet efficiency.
Thus, there remains a need to provide a more efficient, economical, open magnet MRI system having a remote region of substantial field homogeneity.
The present invention provides a MRI magnet system that generates two orthogonal magnetic fields to provide a remote field of homogeneity. The two axis field polarization topology provided by the present invention has magnet system symmetry in two axes providing two orthogonal magnetic fields. For example, a pole pair along the Z-axis generates a remote field in the XY plane and a pole pair along the Y-axis generates a remote field in XZ plane. When using a pole pair along the Z-axis and a pole pair along the Y-axis, i.e., two pole pairs having orthogonal axes, the fields are additive as the vector sum of the two orthogonal fields and provide a maximum in the field strength in the X direction.
The present invention provides a planar MRI system having an open magnet configuration that produces a magnetic field having a substantial remote region of homogeneity. The MRI system includes spatial encoding gradient coils and a rf coil, each preferably having a planar configuration. In a preferred embodiment of the present invention, the magnet configuration comprises, for each of two orthogonal directions, a substantially planar ferromagnetic core around which is wound a pair of primary coils, the coils providing primary current wires on one side of the planar core emitting a flux in the imaging region to generate a resulting magnetic field which provides a remote region of substantial field homogeneity. Preferably, the same planar ferromagnetic core is used for both pairs of primary current wires.
The ferromagnetic core preferably is a plate of a material having high permeability. Further, the core preferably is made by laminating layers of ferromagnetic material to form the substantially planar structure of the core.
The planar ferromagnetic core provides a magnetic shield between the primary current wires on one side of the core plane and the return wires of the current loops on the opposite side of the core plane. In addition, the planar ferromagnetic core is constructed to provide a refractive effect that magnifies the magnetic field provided by the current wires by an imaginary current (xe2x80x9cmirror imagexe2x80x9d) effect when compared to the field provided by the current wires alone. Preferably, the distance between the primary current wires emitting the field and the return wires of the current loops is minimized to reduce the amount of copper and the loss effects for the coil having a particular length of current wire on the surface of the planar core facing the target volume having a region of substantial field homogeneity. In other words the thickness of the planar core preferably is minimized while also providing maximum field strength. However, it also is preferred that the core material adjacent the coil wires be near saturation (but not saturated) when the magnet system is operated. Thus, the core is designed preferably to prevent saturation while operating near saturation in a region near the current wires to obtain a maximum magnetic induction (i.e., B-field) in the core for the ampere turns provided.
In a preferred embodiment of the invention, at opposite ends (i.e., poles) of the planar ferromagnetic core, a wall of ferromagnetic material is provided, for example the core ends can be extended substantially perpendicular to the plane of the core (as defined by the portion of the core lying between the pair of current loops), on the side of the core on which lies the desired remote field of substantial homogeneity. The perpendicular extension of the ends provides an additional magnification effect of the primary current wires to the resulting field strength in the remote region of substantial field homogeneity. More preferably, the ends are extended at an acute angle with the plane of the core lying between the pair of current loops. The angle of the end extensions measured from the perpendicular is about 20xc2x0 or less. Preferably, the angle of the end extensions measured from the perpendicular is about 15xc2x0 or less. Most preferably, the angle of the end extensions measured from the perpendicular is in the range of about 10 to 15xc2x0. The vertical, or angled, end extension or wall structure is provided symmetrically for each of the two axis orthogonal topologies.
Preferably, the end extensions also are made by laminating layers of ferromagnetic material, similar to the construction of the planar ferromagnetic core.
In other preferred embodiments, more than one pair of primary current wires is used for each orthogonal axis magnet topology to provide the desired remote region of substantial field homogeneity. In addition, one or more pair of shimming current wires preferably are used for each orthogonal axis magnet topology to compensate for deflections in the field profile provided by the primary current wires. The shimming current wires preferably can be located on a plane that is closer to the remote region of substantial field homogeneity than the plane of the primary current wires. The use of a shimming planar ferromagnetic core (which typically is much thinner than the primary ferromagnetic core and can be in sections corresponding to the location of particular shimming current wires) also is preferred for the shimming current wires to obtain magnet efficiency.
The magnet configuration of certain embodiments of the present invention can be used to provide two back to back magnet pairs with a common core to provide a two sided MRI system because the ferromagnetic core provides shielding of one field from the other, although producing the two fields on opposite sides of the core with the same primary coils. Thus, a further economy can be obtained where such a back to back configuration can be utilized to provide double patient throughput such as, for example, in mass screening applications such as mammography screening.
Preferably, the magnet configuration providing the remote region of substantial field homogeneity is located on one side of a plane, which is parallel to the plane of the electromagnet ferromagnetic core and separates a patient or body component to be imaged from the magnet configuration, thereby providing a planar open magnet configuration. As described herein, the Y-axis and Z-axis define the orientation of the planar surface, and the X-axis is perpendicular to the planar surface in a rectangular coordinate system. Thus, the background field for the two orthogonal axis magnet system has a direction substantially parallel to the ZY-plane in the rectangular coordinate system and is the vector sum of the first field in a direction parallel to the Z-axis generated by one of the primary current pairs plus the second field in a direction parallel to the Y-axis generated by the other one of the primary current pairs.
As used herein, the term xe2x80x9cremotexe2x80x9d means that the field region is located external to the plane of current wires producing the magnetic field. As used herein, the terms xe2x80x9csubstantial homogeneityxe2x80x9d, xe2x80x9csubstantial field homogeneityxe2x80x9d or xe2x80x9csubstantial relative field homogeneityxe2x80x9d refer to and mean a region having magnetic field homogeneity sufficient for producing MRI images of the desired quality in the region of field maximum.
As used herein, the term xe2x80x9csubstantially planarxe2x80x9d means that the surface of the core facing the desired magnetic field is a flat surface. It should be noted that the thickness of the core for a particular core material will depend upon the desired magnitude of the field to be provided by the magnet, i.e., upon the number of ampere-turns used to excite a given field.
The distance of the region of homogeneity from the planar surface can be controlled basically by varying the spacing between the primary coil pairs. The primary coil pairs are the major source for background field and determine the basic field strength. Preferably using minimal additional power, the shimming coil pairs provide compensation for field deflection resulting from the primary coils and contribute to the degree of homogeneity of the background field. The size, geometry and spacing of the coils can influence the size of the homogeneous region.
The MRI system also preferably has (i) a planar xyz gradient coil system that produces a gradient field for spatial encoding in the region of homogeneity, i.e., the imaging volume or target region, as well as (ii) rf excitation and receiving coil (or coils).
The size of the magnet configuration and MRI system in accord with the present invention can be varied to provide whole body imaging or portable systems for localized imaging.